1. Field of the Invention
The present invention relates to a radiographic image detector and a control method for the radiographic image detector.
2. Description of the Related Art
In medical field, radiography systems utilizing radioactive rays such as x-rays for imaging are known. An x-ray radiography system includes an x-ray projector having an x-ray source for projecting x-rays toward a subject, an object to be imaged, and a radiographic image detector that receives x-rays from the x-ray projector after penetrating the subject, thereby to acquire a radiograph or x-ray image that represents information on the subject. The x-ray source will project x-rays under given image acquisition settings. The image acquisition settings include tube current that decides the amount of x-ray dose per unit time, and tube voltage that decides the quality of x-rays, i.e. the energy spectrum of the x-rays. These image acquisition settings are determined for each test subject according to demands and conditions of the subject to be imaged, such as the target site to be inspected, the age of the subject when it is a patient, etc.
As the radiographic image detector, those using a flat panel detector (FPD) in place of conventional x-ray film or a radiographic imaging plate (IP) have been adopted in practice. As described for example in JPA 2010-264181, the FPD includes a detection panel having a large number of pixels arranged in a matrix in an imaging area. The pixels accumulate signal charges corresponding to the amounts of x-rays incident on the respective pixels, and a signal processing circuit of the FPD reads out the signal charges from the pixels to output the signal charges as an analog voltage signal.
The detection panel includes numbers of pixels or detection elements in an imaging area, each pixel including a photodiode and a thin film transistor (TFT), and a scintillator (a phosphorous member) for converting x-rays to visible light is provided on the imaging area. The TFT is a switching element for switching an electric connection between the photodiode and a signal line on and off to change over operation modes of the pixel. When the TFT is turned off, the photodiode is electrically disconnected from the signal line, so that the photodiode begins to accumulate signal charges. When the TFT is turned on, the photodiode is electrically connected to the signal line, so that the signal charges generated in the photodiode are read out through the TFT and the signal line.
Unlike the conventional x-ray film or imaging plate, the TFTs of the FPD must be turned on or off to start accumulating or reading the signal charges in synchronism with the radiation timing from the x-ray source. To control the FPD synchronously with the x-ray source, a synchronizing signal may be transmitted from the x-ray source to the FPD. Alternatively, as disclosed in JPA 2010-264181, the FPD may automatically detect the start and stop of x-ray radiation based on the intensity of x-rays incident on the FPD, radiation doses per unit time. As the timing of radiation from the x-ray source being detected by the FPD without any communication between the x-ray source and the FPD, the FPD does not need to mount any communication interface for the x-ray source, which allows simplifying the structure of the FPD. Subsequently, any hardware relating to the communication interface, such as cables, terminals, and substrates, may also be omitted. Therefore, the FPD that can detect the radiation timing by itself may also save the cost of the radiographic image detector as the whole.
As illustrated in FIG. 12 of the above prior art, the x-ray intensity begins to increase gradually upon a start command received on the x-ray source, and reaches a peak level that is determined by the tube current. Thereafter, the x-ray intensity is kept at a substantially constant level around the peak level till the x-ray source receives a stop command. Upon the stop command, the x-ray intensity begins to decrease gradually. Therefore, the intensity of x-rays dosed for one image acquisition changes along an approximately trapezoidal time curve.
According to the above prior art, the FPD measures the x-ray intensity as a voltage signal and detects the end of radiation when the voltage gets below a certain threshold level, or when the time derivative of the voltage gets above a certain negative threshold level, i.e. when the gradient of a declivity in a time curve that represents the decreasing x-ray intensity with time gets greater than a predetermined value. Immediately after the end of radiation is detected, the TFTs are turned on to terminate the charge accumulating operation and start the charge reading operation.
However, starting the reading operation at a constant timing from the time of detection of the end of radiation, like in the above prior art, has a problem as set forth below.
Generally the response of the x-ray source to the stop command is so slow that there is a certain time lag between the start of declination of x-ray intensity upon the stop command and the complete stop of radiation, i.e. when the x-ray intensity becomes zero. Hereinafter, the time lag from when the x-ray intensity begins to decrease till it gets down to zero may be referred to as the duration of a radiation wave tail in a time curve representing the course of x-ray intensity along time. The duration of radiation wave tail varies depending upon the tube voltage applied for activating the x-ray source. When the tube voltage is low, the x-ray intensity will decrease more rapidly after the stop command; the radiation wave tail will have a steeper declivity and hence the duration of the radiation wave tail is relatively short. On the other hand, when the tube voltage is high, the x-ray intensity will decrease more slowly after the stop command; the radiation wave tail will have a gentler declivity and hence the duration of the radiation wave tail is relatively long.
As the tube voltage differs from one imaging session to another, the duration of radiation wave tail may vary correspondingly. If the reading operation is predetermined to start in a constant time after an end of radiation detected, the reading start time tends to unexpectedly deviate from the time when the radiation completely stops. For instance, if the time from the detection of end-of-radiation to the start of reading operation is fixedly adjusted to a short radiation wave tail at a low tube voltage, the reading operation will start before the complete end of radiation when the tube voltage is higher and the radiation wave tail lasts longer. To the contrast, if the time from the detection of end-of-radiation to the start of reading operation is fixedly adjusted to a long radiation wave tail at a high tube voltage, the reading operation will start with a certain delay from the complete end of radiation when the tube voltage is lower and the radiation wave tail ends shorter.
Since the pixels are still irradiated with x-rays during the radiation wave tail, indeed the x-ray intensity gradually decreases, if the accumulating operation is terminated to start the reading operation before the complete end of radiation, those x-rays radiated after the end of accumulating operation will not be utilized for image signal acquisition. The waste of radiation may lower the S/N ratio of the image.
Moreover, starting the reading operation before the end of radiation wave tail may cause shading artifact. Shading artifact appears as density gradations increasing in one direction within the image. Since the signal charges are read out from the pixels line after line to get a frame of image, there is a time lag from the reading operation of the first pixel line to the reading operation of the last pixel line. Accordingly, the pixels of the last line can accumulate charges for a longer time than the pixels of the first line. As a result, if the pixels are irradiated with x-rays during the reading operation, the accumulated charges will increase in ascending order of pixel lines, so will the image density distribute. Thus, shading artifact appears in the image, increasing image density gradually in the charge reading direction, i.e. in one column direction orthogonal to the line direction of the pixels.
On the other hand, the photodiodes generate dark currents even while they are not irradiated with x-rays. Therefore, if the accumulating operation is continued after the complete end of radiation, i.e. without any x-ray irradiation, the dark current noise will increase in the image signal, which may also lower the S/N ratio of the consequent image.
For the reasons as above, for the benefit of image S/N ratio improvement, the reading start time should ideally be coincident with the end of radiation wave tail when the x-ray intensity gets down to zero. In order to start the reading operation at the end of radiation wave tail, the threshold level of the x-ray intensity used for detecting the end of radiation may be set as low as substantially zero. However, as described above, because of the risk of detection error due to dark current from the photodiodes, it is hard to adopt such a low threshold level for detecting the end of radiation. The above-mentioned prior art does not disclose nor imply the above problems and solutions therefor.